1. Field of the Invention
The present invention relates in general to gamma ray imaging cameras, and it more particularly relates to such a camera for use in nuclear medicine applications.
2. Background Art
In nuclear medicine, a radioactive isotope is introduced into the area of the body under examination. The radioactive isotope generates gamma rays at certain energies dependent upon the particular isotope used. For example, .sup.99m T.sub.c generates gamma rays having an energy of 140 keV which is used in many nuclear medicine applications.
A collimator is then positioned between the patient and the gamma ray imaging camera or detector. The collimator typically comprises a lead plate having a plurality of parallel throughbores so that the gamma rays which pass through the collimator to the imaging detector are essentially parallel to each other.
A single sodium iodide crystal doped with thallium is contained within the housing for the image detector so that gamma rays passing through the collimator pass through an entrance window on the housing and impinge upon the crystal. A thin aluminum sheet across the entrance window hermetically encloses the crystal to protect it from moisture and the elements while a light reflective surface is positioned between the window and the crystal.
The sodium iodide crystal forms the scintillator material and has a typical thickness of between 0.25-0.50 inches. The crystal in turn is glued to a thick glass sheet known as the light pipe. The light pipe not only transmits photons which are generated by the crystal, but also mechanically supports the crystal. In order to detect the emission of photons from the crystal, an array of photomultiplier tubes (PMT) am optically coupled to the light pipe assembly through an optical coupling compound or cement.
In operation, the generated gamma rays pass through the collimator and are absorbed by the sodium iodide crystal. Only very few gamma rays are absorbed by the aluminum sheet across the entrance window due to the low atomic number of the aluminum. Conversely, the relatively high atomic number of the crystal makes it a good absorber of gamma rays. Most of the gamma rays that are of interest to nuclear medicine fall within the energy range of 60-360 keV. These gamma rays interact with the crystal through the photoelectric effect in which a bound electron in the crystal absorbs the gamma ray. Upon doing so, the entire energy of the gamma ray is transferred into kinetic energy of the electron.
The electron which absorbs the energy of the incoming gamma ray transfers this energy to adjacent electrons through coulomb collisions. Many electron-hole pairs are formed in the crystal as a result of the energy deposited by each gamma ray photon. The recombination of the holes in the electrons then creates a large number of low energy scintillation photons. Measurements have shown that for a sodium iodide scintillation crystal, 11.4-13.5% of the total energy from the absorbed energy is emitted as scintillation photons.
The scintillation photons are emitted in random directions by the crystal. Those photons striking the light reflective surface are reflected back towards the photomultiplier tubes so that for each gamma ray absorption, photons are emitted in a conical pattern towards the photomultiplier tubes with a higher concentration of the photons in the center of the cone. These photons are detected by the photomultiplier tubes and, for best possible resolution of the gamma ray absorption, the photons should strike at least seven photomultiplier tubes. Well known electronic circuitry is then employed to determine the position of the gamma ray absorption from the output signals of the photomultiplier tubes.
The scintillation photons travel radially outward from the point of absorption in random directions. The photons travelling towards the photomultiplier tubes are refracted at the crystal/light pipe interface due to a mismatch in the refractive indices for sodium iodide (N=1.85) and glass (N=1.5). Not all of the photons which strike the photomultiplier tubes generate a photoelectron from the photocathode. Instead, the quantum efficiency of the photomultiplier tube is expressed as a percentage, i.e. the percentage of the number of photons striking the photomultiplier tubes which generate a photoelectron from the photocathode.
These previously known gamma ray imaging detectors, however, suffer from a number of disadvantages. One main disadvantage is that the photomultipliers are bulky and heavy, rendering the gamma ray cameras impractical to move around from one location to anther, and to be readily available. Yet a further disadvantage is that the sodium iodide crystal generates only a relatively small number of scintillation photons per absorbed gamma ray. Another disadvantage is that the sodium iodide crystals must be hermetically sealed by the manufacturer in order to protect the crystal from humidity absorption. The sodium crystals are also brittle and can be easily fractured by temperature or thermal shock.
One attempt to resolve some of these problems is described in Engdahl et al., U.S. Pat. No. 5,171,998, which discloses a gamma ray imaging detector having a single scintillation detector crystal which converts absorbed gamma rays into a plurality of scintillation photons. The Engdahl U.S. Pat. No. 5,171,998 patent is incorporated herein by reference. A cesium iodide crystal doped with thallium [CsI(T1)crystal] in place of the sodium iodide crystal as the scintillating detector crystal.
An array of photodiodes is arranged along one side of the crystal to receive the scintillation photons which generate an electrical output signal proportional to the number of scintillation photons received by the photodiode. Diodes with low capacitance and electrical noise, such as silicon drift photodiodes, are employed so that the signal generated by the photodiode as a result of the received scintillation photons is readily detectable above the electrical noise from the photodiodes. An electronic circuitry is then utilized to determine the position of impingement and absorption of the gamma ray.
The PCT patent application. WO93/01612 to Nudelman describes a large area video camera suitable for high energy imaging applications. The sensor-target of the camera tube is composed of T1Br or a two layer structure comprising CsI and a photoconductive layer of materials such as amorphous silicon, amorphous selenium, cadmium sulphide, antimony trisulphide or antimony sulphide oxysulphide. The disclosed tube further deals with problems associated with stray capacitance.
More particularly, this patent describes a low velocity electron beam photoconductive-type video tube which employs novel sensor-target configurations and incorporates a modified electron optical system to acquire large images defined by penetrated ionizing radiation such as X-rays and gamma rays. It is suitable for applications in nuclear medicine, diagnostic radiology and non-destructive testing. In one embodiment, a large area X-ray sensitive video camera tube has a sensor-target including a signal plate for detecting irradiating photons and for providing sufficient storage capacity to hold the electrons on the surface of the signal plate.
The target includes a scintillator which comprises a layer substantially composed of CsI. The CsI layer generates a visible light photon output proportional to the absorbed radiation. A photoconductor which is responsive to the light photon output forms an electronic image of the radiation and comprises a substrate substantially composed of a material selected from the group consisting of amorphous selenium, antimony trisulphide, cadmium sulphide and antimony sulphide oxysulphide. The electron optics of the tube generates a low velocity electron beam which is directed in raster fashion at the photoconductor so that an electronic image from high energy radiation directed at the target is acquired at the signal plate for transmittal to the video readout circuit. This combination comprising a layer of CsI plus an adjacent photoconductive layer is designated as the Cs+ sensor-target.
The CsI layer may be doped with Na or Tl. The CsI layer has a thickness that depends upon the X-ray or gamma ray energy requirement of the application. The large area imaging system may further include an image processor and a film writer to provide a hard copy readout of the electronic image. The tube includes electron optics for generating a low velocity electron beam and directing the beam in raster fashion at the sensor target. The electron beam in tracing out a raster deposits a uniform surface charge of electrons. The sensor-target absorbs the high energy radiation of the imaging beam, causes electrons to be removed and results in a new charge distribution which is an electronic reproduction of the X-ray photon distribution, i.e., the intrinsic X-ray image. The CsI in the scintillator layer may be doped with Na to provide a predominantly blue light emission or alternately Tl which produces a predominantly green light emission.
While the disclosed gamma ray imaging detectors theoretically resolve some of the concerns of conventional devices, such as the elimination of photomultipliers, they do not resolve other significant disadvantages such as the individual recordation of the gamma rays. None of the above patents or other publications disclose a gamma ray camera which is small in size, light in weight, easily transportable, and which integrates the recordation of the gamma rays over the total number of events, rather than individually recording each gamma ray separately.